Ultrasound is a term that refers to acoustic waves having a frequency above the upper limit of the human audible range (i.e., above 20 kHz). Because of their relatively short wavelength, ultrasound waves are able to penetrate into the human body. Based on this property, ultrasound in the frequency range of 2–20 MHz has been widely used to image internal human organs for diagnostic purposes.
To avoid thermal damage to tissue, the power level in diagnostic ultrasound imaging is kept very low. The typical ultrasound intensity (power per unit area) used in imaging is less than 0.1 watt per square centimeter. High intensity focused ultrasound, which can have an intensity above 1000 watts per square centimeter, can raise the tissue temperature at the region of the spatial focus to above 60 degrees Celsius in a few seconds and can cause tissue necrosis almost instantaneously.
High intensity ultrasound has been proposed to treat and destroy tissues in the liver (G. ter Haar, “Ultrasound Focal Beam Surgery,” Ultrasound in Medicine and Biology, Vol. 21, No. 9, pp. 1089–1100, 1995); in the prostate (N. T. Sanghvi and R. H. Hawes, “High-intensity Focused Ultrasound,” Experimental and Investigational Endoscopy, Vol. 4, No. 2, pp. 383–395, 1994); and in other organs.
Ultrasound transducers generate ultrasound waves for imaging and therapy. A typical ultrasound transducer comprises piezoelectric materials such as PZT ceramics, electrodes, matching layers, and backing materials. When an electrical field is applied to two electrodes on the opposite sides of a piezoelectric ceramic plate, the thickness of the plate expands or contracts, depending on the polarity of the field. If the electrical field polarity alternates at a high frequency above 20 kHz, the mechanical vibration caused by the rapid expansion/contraction of the plate generates ultrasound waves.
During ultrasound therapy, high electrical power is applied to the ultrasound transducer to generate a correspondingly high acoustical output power. Transducer power conversion efficiency is the ratio of the output acoustic power to the input electrical power. A high transducer power conversion efficiency is always desirable to minimize the transducer internal heating due to electrical power losses.
During ultrasound imaging, low-power electrical pulses drive the transducer, causing it to transmit the low power ultrasound pulses into the patient body. Ultrasound echoes, reflected from organ boundaries and other tissue and physiological structures within the body, are typically received by the same ultrasound transducer and converted to electrical output signals, which are processed to produce ultrasound images of the internal organ on a display. A transducer having a broad frequency bandwidth is desirable to obtain good image resolution. Often, however, the desire for high efficiency during ultrasound therapy and the desire for broad bandwidth during ultrasound imaging are difficult to satisfy simultaneously in the same transducer design.
To treat or to image a large volume of diseased tissue, the ultrasound beam is caused to scan through the tissue, either mechanically or electronically. In a mechanical scanning device, such as disclosed in U.S. Pat. No. 4,938,216, one or more electrical motors position the ultrasound transducer in different positions. One of the more common types of electronic scanning device employs an ultrasound linear phased-array transducer, such as that disclosed by E. B. Hutchinson and K. Hynynen, in an article entitled “Intracavitary Ultrasound Phased Arrays for Noninvasive Prostate Surgery” (IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 43, No. 6, pp. 1032–1042, (1996)), and in U.S. Pat. No. 4,938,217. An electronic scanning device has a plurality of small piezoelectric elements disposed in an array. These elements are independently driven. By properly controlling the phase of the driving signals applied to energize these elements, the array will be caused to form ultrasound beams directed at different depths and angles. The electronic scanning transducer has many advantages over the mechanically scanned transducer. The main advantage is that there are no moving components in the electronic device, so that it has much higher durability and reliability. The disadvantage of the electronic device is its complexity and associated relatively high cost. To achieve a compromise between the advantages and the disadvantages, some prior art references, such as U.S. Pat. No. 4,757,820, disclose transducer designs that include both the mechanical and the electronic approaches.
However, system and transducer complexity is still one of the major disadvantages of electronic therapeutic arrays. A therapeutic transducer requires a large surface area to generate a high acoustic power output and a large aperture for deep treatment. Preferably, the f-number (focal depth over aperture size) is kept constant within the range from 0.8 to 2.5. On the other hand, to steer the ultrasound beam over a wide range and to focus the beam using a small f-number, the ultrasound phased array must have very fine, narrow array elements, because a narrow element can transmit an ultrasound beam throughout a wide range of directions.
To provide a transducer having both a large aperture and fine elements to enable it to provide imaging and therapy functions, a conventional therapeutic phased array design includes a very large number of elements. For example, to treat lesions at a maximum depth of 5 cm, a therapeutic linear array having an f-number of 1.0 should have an aperture width of about 5 cm. For use at this depth, the transducer will typically operate at a frequency of about 3 MHz. The wavelength of ultrasound in water or biological soft tissue at this frequency is about 0.5 mm. For a phased array of this configuration to have a sharp focus (i.e., a relatively small f-number), the array typically will have an element pitch size about 0.5 to 0.7 times the wavelength of the ultrasound beam it produces. For a pitch size of about 0.6 times the wavelength, an exemplary therapeutic array might have an element pitch size of about 0.3 mm and a total of about 167 elements.
Each element has a dedicated electronic driving circuit in a control system for the array. To drive a phased array like that discussed above, the control system would need to include 167 sets of driving circuits, i.e., one for each element. The array and the control system are connected through a thick cable that includes at least 167 smaller coaxial cables inside it. Each smaller coaxial cable should have conductors of a sufficiently large cross-sectional area to carry a relatively large current to the therapeutic array element. The thick cable required to meet this need makes the device difficult to handle.
Considering all these constraints, it will be evident that the complexity of such a therapeutic phased array, including the cable and the control system coupled to it, can easily become impractical to engineer, and its cost will most certainly exceed the budget of most medical facilities. It is for these reasons that the therapeutic phased array has not been widely accepted.
It would be desirable to use an ultrasound array transducer for both imaging and therapy. The smaller size of a probe having a transducer that is usable for both functions is an advantage. For example, in many endoscopic, therapeutic-ultrasound applications, there are limitations on the size of the treatment devices that can be employed. Thus, a dual-purpose ultrasound array transducer may save space in the probe. Also, in ultrasound image-guided therapeutic applications, there are two spatial planes, one for imaging and the other for treatment. These two planes should overlap so that the treatment area can be observed in the imaging plane. Oftentimes, however, it is difficult to register the two planes from two spaced-apart transducers. Sometimes, there are blind spots in the treatment zone, which are not observable in the imaging plane. However, if one transducer is used both for imaging and treatment, the problem of non-overlapping zones does not arise.
The prior art has not dealt extensively with the problem of designing a dual-purpose phased array transducer. Besides the conflict between the disparate design parameters that must be satisfied to achieve efficiency and adequate bandwidth in such a transducer, as noted above, there are other unresolved issues in making a therapeutic phased array transducer, such as heat dissipation, and element cross-talk. In U.S. Pat. No. 6,050,943 and in an article published by P. G. Barthe and M. H. Slayton, entitled “Efficient Wideband Linear Arrays for Imaging and Therapy” (IEEE Symposium in Ultrasonics, Ferroelectrics and Frequency Control, November 1999), the authors address some of these problems.
Thus, there is a clear need for an ultrasound device that employs simple and highly efficient ultrasound transducer arrays usable for both imaging and therapy. This kind of ultrasound device can be used to generate real-time ultrasound images of a patient's internal condition, provide ultrasound therapy to a treatment site, and monitor the treatment results. Such an ultrasound transducer should have variable geometry for treating different pathologies. In addition, the transducer array should be capable of generating high-intensity ultrasound to ablate or necrose tumors and other diseased tissues.